The use of implantable electrodes has long been known in the medical arts. Electrodes are used, for example, to deliver electrical stimulation to cardiac tissue for pacing, cardioversion, and defibrillation applications. Implanted electrodes are also used to stimulate nerve tissue to modulate cardiac activity, and to treat other cardiovascular disorders. More recently, implanted electrodes have been used to treat incontinence, gastro-intestinal problems, and neurological disorders. Other types of ailments and physiological disorders are treated using similar electrical stimulation.
One problem associated with the use of implanted electrodes to deliver electrical stimulation to living biological tissue involves the accumulation at the electrode/tissue interface of a double layer of charge. More specifically, when an electrode is placed in contact with biological material such as tissue, a layer of electrons accumulates at the conducting surface of the electrode. In response, a corresponding layer of positively-charged ions accumulates within the biological matter near the electrode surface. This double-layer of charge at the molecular level imports a capacitance into the equivalent circuit. This capacitance, sometimes referred to as capacitive impedance, has an effect not unlike an ordinary parallel-plate capacitor, acting as a high-pass filter that distorts signals recorded via the electrode. The impedance minimally affects high-frequency signals, but imposes significant, non-linear attenuation at lower frequencies. Moreover, this capacitance imports a phase shift that varies with frequency. Biological signal recording is significantly impacted since physiological signals include frequency components of 100 Hz or less. Although the foregoing example utilizes a parallel-plate capacitor to illustrate the attenuation imposed at the electrode/tissue interface, this analogy is not entirely accurate. The capacitive impedance Z associated with an ordinary parallel-plate capacitor is inversely proportional to the signal frequency. This relationship can be expressed as
 |Z|=1/(□C)
wherein □ is the angular frequency that correlates to signal frequency f via the equation □=2□f. C is a constant referred to as the capacitance, which has a value that is dictated by the geometry, material construction, and potential difference appearing across a given capacitor.
In contradistinction to parallel-plate capacitors, the capacitive impedance at an electrode/tissue interface is inversely proportional to the square root of the signal frequency. This may be expressed as|Z|=k(1/□□)where k is a constant, and □ is the angular frequency of the signal. This relationship is different from that discussed above with respect to parallel-plate capacitors because of the ionic diffusion that occurs at tissue/electrode interfaces. In other words, the transfer of a signal across an electrode/tissue interface involves the movement of ions within the tissue surrounding the electrode. In contrast to a signal that is transferred across a parallel-plate capacitor via movement of electrons e− and electron vacancies h+, this ionic movement occurs more slowly, resulting in a larger impedance.
All materials that are currently employed by implantable electrode systems possess a capacitive element such as described above. For instance, all noble metal electrodes and all metal electrodes having a thin passivating film coating, exhibit the type of impedance associated with ionic diffusion. A partial list of such materials includes platinum (Pt), platinum-iridium alloy/s (Pt—Ir), platinized platinum, gold (Au), titanium (Ti), titanium nitride (the nonstoichiometric interstitial nitride of titanium, TiNx where x varies from 0.8 to 1.15), stainless steel, silver-silver chloride (Ag|AgCl or Ag/AgCl), iridium oxide, alloys of metals of all compositions and components, as well as carbon, glassy carbon and vitreous carbon.
In traditional parallel plate capacitors, capacitive impedance may be decreased by increasing the capacitor surface area. This increased surface area proportionally increases the value of constant C so that the impedance |Z| decreases to zero for a larger and larger surface area. In a similar manner, increasing an electrode surface reduces constant k, reducing the impedance |Z| at the electrode/tissue interface. Although increased electrode area, may reduce impedance, the signal distortion is not completely avoided. Moreover, this increased electrode surface area may be undesirable for several reasons. For example, the increased surface area may result in a lower current density, thereby necessitating the increase of stimulation parameters such as pacing threshold levels, which, in turn, negatively impacts battery life. Moreover, the larger electrode assembly may be more difficult to deliver to a target destination.
The foregoing discussion addresses the problems associated with capacitive impedance at an electrode/tissue interface. In addition to capacitive impedance, other types of galvanic and Faradaic impedances may cause distortion of biological signals measured by implanted electrodes. For example, redox reactions may occur between an electrode surface and ambient chemical species that are not intrinsic to the function of the electrode. These reactions, which may be modeled by amplifiers, diodes, and other non-linear attenuating circuit elements, increase the distortion of recorded signals by importing unpredictable, time-dependent negative impedances.
What is needed, therefore, is an implantable electrode system that provides reduced signal distortion and enhanced signal recovery. Ideally, the resulting electrode/tissue interface is ohmic at physiological frequencies.